This invention refers to a device capable of rapidly and efficiently converting the energy in a laser pulse into a high-impulse pressure wave and to its applications in the transient permeabilization of a biological barrier, including the outer layers of the skin, without causing damage or discomfort. The invention is described in terms of the physical principles of efficient laser generation of pressure waves, of the properties of the absorbing materials that maximize the amplitude of the pressure wave and minimize its rise time, and of the practical delivery of drugs or biologically active compounds to or through the skin or other biological barriers. Examples are given to illustrate the use of the device and its applications. The distinctive features of the device are the use of thin films (thickness <200 μm) strongly absorbing the light of pulsed lasers and the use of affordable pulsed lasers, such as lasers with low energies (laser energy <500 mJ), wherein the said device is capable of generating high impulse acoustic transients with short rise times (rise time <50 ns) at low optical power densities (<40 MW/cm2 per pulse).
The skin is a very effective protection against the ingress of foreign material, such as chemicals and microbes. The outermost layer of the skin, called stratum corneum (SC), is only 10 to 20 μm thick but it is the main contributor to the skin's impermeability. It is made by a dozen layers of hardly packed nonliving corneocyte cells, embedded in a mixture of lipids with high spatial organization. Most molecules penetrate the skin by diffusion through the intercellular lipids, a tortuous path around the corneocytes that is highly constrained by structural and solubility requirements. Underneath the SC there is the viable epidermis, which is 50 to 100 μm thick and is vascular. Further below is the dermis, 1-2 mm thick and rich in capillaries, capable of clearing most penetrants within minutes [1].
The idea of delivering drugs through the skin is centuries old. The attempts to attain this goal can be classified in two groups: passive and active drug delivery [2]. The first class refers to formulations of vehicles optimized to enhance the diffusion of a particular drug through the skin, including ointments, creams or gels, which may include chemical permeation enhancers. Indeed, a wide variety of chemicals have the ability to increase skin permeability, such as dimethyl sulfoxide, laurocapram (Azone), 2-n-nonyl-1,3-dioxolane (SEPA), fatty acids and fatty acid esters, surfactants, and others well known in the art [3]. However, the increased permeation enhancement, even of small molecules, typically correlates with increased skin irritation [4]. Passive methods are only efficient for transdermal delivery of small molecules (molecular weight <500 Da) with adequate lipophilicity (n-octanol-water partition coefficients KOW, in the range 1<(log KOW)<3) and with less than 3 hydrogen-bonding groups. Moreover, the formulation of the vehicle is specific for a given drug. Active drug delivery methods employ physical methods such as electrical assistance (iontophoresis, electroporation), mechanical processes (microneedles, abrasion, ablation, perforation, microprojections), or ultrasounds (sonophoresis). Another active method of drug delivery consists in generating photomechanical waves by intense pulsed-laser irradiation of a target [5]. The stress waves formed by optical breakdown, ablation or thermoelastic expansion have been shown to transiently permeabilize the SC and facilitate the transport of macromolecules into the viable epidermis [6].
Technical applications of stress waves in the delivery of compounds through epithelial cell layer have been described. For example, Kollias et al [7] described in U.S. Pat. No. 6,251,099 B1 a compound delivery using impulse transients generated by lasers with fluences between 1 and 7 J/cm2 and pulse widths of 20-30 ns, which correspond to optical power densities between 40 and 300 MW/cm2. These intense and short laser pulses were directed to targets with thickness ranging from 0.8 mm (for metals) to 3 mm (for plastics). At these optical power densities, the dominant mechanism of interaction between the laser pulse and the target is ablation of the target material and is usually accompanied by plasma formation. This produces ejection of material from the surface of the target hit by the laser pulse and the associated recoil momentum propagates in the bulk of the target to reach its opposite surface as an acoustic wave. The ability of this acoustic wave to increase temporarily the permeability of the skin has been related to its impulse. Intradermal delivery of large compounds without damaging the skin requires impulses between 2 and 50 bar/ns [5]. The generation of pressure waves with such high impulses in a useful area of a current target requires lasers with high power densities that are recognized as complex and costly, and alternatives have been sought [8].
Apparatus for enhancing drug delivery using optical power densities down to 10 mJ/cm2 were also described. For example, Visuri et al [9] described in U.S. Pat. No. 6,484,052 B1 how such low power densities at laser pulse frequencies between 100 Hz and 1 MHz can be coupled to a fiber optic and inserted in a portion of a human body to generate an acoustic radiation field in that portion. They also described the attachment of an optically-powered mechanical transducer to the distal end of said fiber, but failed to specify the characteristics of such a transducer. Unless the transducer has very specific physical, photochemical and material properties, it will not be able to produce a pressure wave capable of transiently permeabilizing biological barriers. The use of high laser pulse frequencies, rather than a single or a small number of laser pulses, does not change the properties of the pressure waves. Hence, no optically-powered mechanical transducer capable of producing acoustic waves capable of permeabilizing biological barriers was disclosed.
In another field, that of biological tissue spectral characterization, Biagi et al described in U.S. Pat. No. 6,519,376 B2 an opto-acoustic generator to generate acoustic, or ultrasound, waves from a pulsed laser-energy source [10]. The absorption of a laser pulse by a graphite-containing layer applied to the tip of an optical fiber connected to the laser source was shown to efficiently produce very-wide-band acoustic pulses. Pure graphite films can also be produced with sub-micrometer thicknesses and still have sufficient mechanical resistance to be handled in pieces of 2.5 cm in diameter [11]. Moreover, with a thickness of 50 nm, the apparent absorptance of these films approaches unity at 400 nm. Free-standing graphite films with very short-lived excited states [12] can be conveniently prepared by pyrolysis of polyacrylonitrile. However, acoustic pulses generated by this method with graphite-containing layers were never considered for drug delivery because the field of tissue characterization employs microjoule laser pulses [13] and this is insufficient for drug delivery with therapeutic effects. Actually, therapeutic effects are not desirable in the spectral characterization, or diagnostic, of biological tissues.
The plethora of transdermal delivery systems does not obscure the fact that they remain a minor alternative to oral delivery or hypodermic injections. A simple and economic transdermal delivery method, capable of delivering a wide variety of drugs through the skin without causing pain or discomfort, and that allows the skin to recover its protective function a few minutes after the application, would confer to transdermal delivery the same status in medical practice as oral delivery or hypodermic injections.